Tissue-engineered constructs

ABSTRACT

The present invention provides constructs including a tubular biodegradable polyglycolic acid scaffold, wherein the scaffold may be coated with extracellular matrix proteins and substantially acellular. The constructs can be utilized as an arteriovenous graft, a coronary graft, a peripheral artery bypass conduit, or a urinary conduit. The present invention also provides methods of producing such constructs.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.13/978,422, filed Jul. 5, 2013. U.S. application Ser. No. 13/978,422 isa United States national phase application under 35 U.S.C 371 ofInternational Application No. PCT/US2012/020513, filed Jan. 6, 2012 andpublished Jul. 12, 2012 as WO/2012/094611. PCT/US2012/020513 claims thebenefit of U.S. Provisional Application No. 61/430,381, filed Jan. 6,2011. The contents of the aforementioned patent applications areincorporated herein by reference in their entireties.

BACKGROUND OF THE INVENTION

There is a considerable need for vascular grafts when the patient's ownvasculature is either unavailable because of prior harvest or unsuitablesecondary to disease. Instances when a vascular graft might be neededinclude peripheral arterial disease, coronary artery disease, andhemodialysis access for patients with end stage renal disease. To date,the most successful vascular conduit for coronary or peripheral vascularsurgery is the patient's own blood vessel, obtained from elsewhere inthe body, often the greater saphenous vein in the leg. For patientsrequiring hemodialysis, the ideal access is a fistula, or a connectionbetween the patient's own artery and vein.

When autologous vessels are not available, syntheticpolytetrafluoroethylene (PTFE) grafts are often utilized for largediameter (≧6 mm) applications, such as arteriovenous access forhemodialysis (U.S. Renal Data System, “USRDS 2009 Annual Data Report:Atlas of Chronic Kidney Disease and End-Stage Renal Disease in theUnited States” (National Institutes of Health, National Institute ofDiabetes and Digestive and Kidney Diseases, 2009) or above the kneeperipheral arterial bypass. However, arteriovenous PTFE grafts forhemodialysis have a poor median patency of only 10 months because ofinfection, thrombus, or intimal hyperplasia-induced occlusion at eitherthe distal anastomosis or outflow vein (U.S. Renal Data System; Schild,et al., J Vasc Access 9, 231-235 (2008)). Other types of grafts, such asdecellularized bovine internal jugular xenografts and human allograftvessels from cadavers, are prone to aneurysm, calcification, andthrombosis, and therefore have not gained widespread clinical acceptance(Sharp et al., Eur J Vasc Endovasc Surg 27, 42-44 (2004); Dohmen et al.,Tex Heart Inst J 30, 146-148 (2003); Madden et al., Ann Vasc Surg 19,686-691 (2005)). In situations where small diameter (i.e., 3-4 mm)vessels are required, such as below the knee and coronary artery bypassgrafting, the patient's own vasculature (i.e., internal mammary artery,saphenous vein) is predominantly used because synthetic grafts andallografts have unacceptably low patency rates (e.g., patency is <25% at3 years using synthetic and cryopreserved grafts in peripheral andcoronary bypass surgeries, compared to >70% for autologous vascularconduits) (Chard, et al., J Thorac Cardiovasc Surg 94, 132-134 (1987);Albers, et al., Eur J Vasc Endovasc Surg 28, 462-472 (2004); Laub, etal., Ann Thorac Surg 54, 826-831 (1992); Collins, et al., Circulation117, 2859-2864 (2008); Harris et al., J Vasc Surg 33, 528-532 (2001);Albers, et al., J Vasc Surg 43, 498-503 (2006)). Thus, a readilyavailable, versatile vascular graft with good patency that resistsdilatation, calcification, and intimal hyperplasia would fill asubstantial and growing clinical need.

To date, tissue engineered vascular grafts formed by seeding autologousbone marrow cells onto a copolymer of L-lactide and ε-caprolactone(Shin'oka, et al., J Thorac Cardiovasc Surg 129, 1330-1338 (2005)), orculturing autologous fibroblasts and endothelial cells (ECs) without ascaffold (McAllister, et al., Lancet 373, 1440-1446 (2009)), have shownpromising functional results in early clinical trials. Thus far, onlythe latter has proven physically strong enough for use in the arterialcirculation. This patient-specific graft requires a 6-9 month cultureperiod in which the autologous fibroblasts produce sheets of tissue. Thesheets are fused together around a stainless steel mandrel (4.8 mmdiameter), inner fused layers are dehydrated, and the graft lumen isseeded with autologous ECs (McAllister, et al., Lancet 373, 1440-1446(2009)). Because of high production costs (≧$15,000 per graft(McAllister, et al., Regen Med 3, 925-937 (2008)) and long wait time (upto 9 months) for patients that require expeditious intervention, it isunlikely that this approach will become standard clinical practice.

Thus, there is a need in the art for effective, rapidly available,reliable and cost-effective tissue engineered constructs that canfunction long term, with minimal to no side effects, in vivo.

SUMMARY OF THE INVENTION

The present invention provides a construct including a tubularnon-woven, biodegradable polyglycolic acid scaffold, wherein the densityof the polyglycolic acid is about 45 mg/cc to about 75 mg/cc and saiddensity is uniform across the entire tubular scaffold.

The length of the tubular biodegradable polyglycolic acid scaffold canbe about 1 cm to about 100 cm. Preferably, the length of the tubularbiodegradable polyglycolic acid scaffold can be about 10 cm to about 40cm. The inner diameter of the tubular biodegradable polyglycolic acidscaffold can be greater than about 3 mm. Preferably, the inner diameterof the tubular biodegradable polyglycolic acid scaffold can be about 3mm to about 20 mm.

The thickness of the polyglycolic acid can be about 0.8 to about 1.5 mmand said thickness is uniform across the tubular scaffold. Preferably,the polyglycolic acid can be about 0.8 to about 1.2 mm and saidthickness is uniform across the tubular scaffold. The thickness of thefibers within the polyglycolic acid can be about 5 to about 20 μm. Theporosity of the polyglycolic acid can be about 90% to about 98%.

The constructs of the present invention can further includenon-biodegradable polyethylene terephthalate supports at each end of thetubular biodegradable polyglycolic acid scaffold. The non-biodegradablepolyethylene terephthalate supports can be attached to the tubularbiodegradable polyglycolic acid scaffold by any means known in the art.Preferably, the polyethylene terephthalate supports are attached viasutures. The porosity of the polyethylene terephthalate can be ≧200cc/min/cm². The tubular biodegradable polyglycolic acid scaffold and thenon-biodegradable polyethylene terephthalate supports can permit theattachment and growth of cells. In other embodiments, othernon-biodegradable polymers can be used to support each end of thetubular scaffold.

The constructs of the present invention are substantially free of heavymetal contaminants. Preferably, the construct includes only traceamounts of heavy metal contaminants selected from the group consistingof: aluminum, barium, calcium, iodine, lanthanum, magnesium, nickel,potassium and zinc.

The constructs of the present invention can further includeextracellular matrix proteins within, and around, the biodegradablepolyglycolic acid scaffold. Preferably, the thickness of theextracellular matrix proteins is greater than about 200 micrometers atthe thinnest portion of the matrix.

The present invention also provides methods of producing a tubularpolyglycolic acid construct including (a) providing a biodegradablepolyglycolic acid sheet, wherein the density of the polyglycolic acid isabout 45 mg/cc to about 75 mg/cc and the thickness of the polyglycolicacid sheet is about 0.8 to about 1.2 mm, (b) wrapping the polyglycolicacid sheet around a mandrel such that opposite edges of the polyglycolicacid sheet meet at an interface; (c) pulling polyglycolic acid fibersfrom each opposing edge of the sheet across the interface, and (d)forming a seam by entangling said pulled polyglycolic acid fibers fromone side of the interface with the polyglycolic acid fibers on theopposite side of the interface, wherein the density of the polyglycolicacid at the seam is about 45 mg/cc to about 75 mg/cc and the thicknessof the polyglycolic acid at the seam is about 0.8 to about 1.5 mm,thereby producing a tubular biodegradable polyglycolic acid constructwith a uniform polyglycolic acid density. The present invention alsoprovides a tubular biodegradable polyglycolic acid construct formed bythe methods described herein.

The construct can be selected from the group consisting of anarteriovenous graft, a coronary graft, peripheral artery bypass conduit,fallopian tube replacement, and a urinary conduit. The mandrel cancomprise any material known in the art. Preferably, the mandrelcomprises a gas permeable, silicone tube.

The entangling step may be performed by any method known in the artwhich permits the seam to remain intact in subsequent treatment steps.Preferably, entangling is performed with a felting needle.

The methods of the present invention can further include, treating thetubular construct to remove heavy metal contaminants. Preferably, thetubular construct is treated with one or more non-polar solventsfollowed by treatment with a primary alcohol, such as ethanol.Preferably, the seam remains intact following said treatment. Thistreatment may also be performed on the biodegradable scaffold prior toformation of a tube.

The methods of the present invention can further include, treating thetubular construct to increase the rate of polyglycolic acid degradationand/or increase the wettability of the polyglycolic acid. Preferably,the tubular construct is treated with a strong base. More preferably,the strong base is 1M NaOH. Preferably, the seam remains intactfollowing said treatment. This treatment may also be performed on thebiodegradable scaffold prior to formation of a tube.

The methods of the present invention can further include, attachingnon-biodegradable polyethylene terephthalate supports at end of thetubular biodegradable polyglycolic acid scaffold.

The present invention also provides a tubular construct comprisingextracellular matrix proteins and polyglycolic acid having an internaldiameter of ≧3 mm, wherein the construct is immune and calcificationresistant, wherein the polyglycolic acid comprises less than 33% of thecross-sectional area of said construct and wherein the construct issubstantially acellular comprising less than 5% cells, less than 2%cells, less than 1% cells or contains no cells. Preferably, the cellsare intact cells. Preferably, the polyglycolic acid comprises less than10% of the cross-sectional area of the construct. More preferably, thepolyglycolic acid comprises less than 5% of the cross-sectional area ofthe construct. Most preferably, the polyglycolic acid comprises lessthan 3% of the cross-sectional area of the construct.

The extracellular matrix protein construct can comprise a burst pressureof greater than 2000 mm Hg. The construct can comprise a suture strengthof greater than 120 g. The inner diameter of the tubular construct canbe about 3 mm to about 20 mm. The thickness of the tubular construct canbe greater than about 200 micrometers at the thinnest portion of theconstruct. The construct can be impermeable to fluid. Preferably, theconstruct is impermeable to fluid leakage up to at least 200 mm Hg, atleast 300 mm Hg, or at least 400 mm Hg. The length of the construct isabout 1 cm to about 100 cm. Preferably, the length of the construct isabout 10 cm to about 40 cm.

The extracellular matrix proteins can comprise hydroxyproline,vitronectin, fibronectin and collagen type I, collagen type III,collagen type IV, collagen VI, collagen XI, collagen XII, fibrillin I,tenascin, decorin, byglycan, versican and asporin. Preferably, theextracellular matrix proteins can comprise hydroxyproline at >40 μg/mgdry weight. In some embodiment, the construct does not comprise elastin,MAGP1 and/or MAGP2. Preferably, the extracellular matrix proteins areproduced from allogeneic, autologous or xenogeneic cells to the intendedrecipient of the construct. Preferably, the extracellular matrixproteins are, in part, oriented circumferentially around the tubularconstruct.

The construct can comprise less than 300 ng/cm of beta-actin. Theconstruct can comprise less than 3% dry weight of lipids. The constructcan comprise trace amounts or no detectable amounts of double strandedgenomic DNA. Preferably, the amount of DNA is as determined by gelelectrophoresis.

The construct induces little to no calcification upon implantation invivo. Preferably, the construct induces less than 1% calcificationwithin 6 months of implantation. More preferably, the construct inducesless than 1% calcification within 12 months of implantation. Mostpreferably, the construct produces no calcification within 12 months ofimplantation.

The construct induces little to no immune response upon implantation invivo. Preferably, when implanted as a vascular graft, the constructinduces less than 1 mm of intimal hyperplasia thickening in nativevasculature and in the graft at anastomoses with the construct at 6months of implantation. More preferably, the construct induces less than0.25 mm of intimal hyperplasia thickening in native vasculature atanastomoses with the construct at 6 months of implantation.

The construct does not dilate greater than 50% beyond its implantdiameter after implantation. The construct may be stored at about 2° toabout 30° C. Preferably, storage at about 2° to about 30° C. istolerated for at least 3 months. Most preferably, storage at about 2° toabout 30° C. is tolerated for at least 12 months.

The present invention also provides methods of producing a tubularconstruct comprising (a) providing a tubular biodegradable polyglycolicacid construct, (b) seeding human cells at passage 6 or less on thetubular biodegradable polyglycolic acid construct, (c) culturing thecells under conditions such that the cells secrete extracellular matrixproteins on the tubular biodegradable polyglycolic acid construct, (d)decellularizing the construct in step (c) such that the construct issubstantially acellular comprising less than 5% cells and wherein theconstruct is immune- and calcification-resistant, and (e) degrading thepolyglycolic acid construct in step (c) such that the polyglycolic acidcomprises less than 33% of the cross-sectional area of said construct,thereby producing a decellularized tubular construct. The presentinvention also provides a decellularized tubular construct formed by themethods described herein.

Preferably, the construct is substantially acellular comprising lessthan 2% cells, less than 1% cells or contains no cells. Preferably, thecells are intact cells. The cells can be allogeneic, autologous orxenogeneic to the intended recipient. Preferably the cells areallogeneic.

The cells are obtained from a single donor or obtained from a cell bank,wherein the cells in the cell bank are pooled from a plurality ofdonors. Preferably, the cells are obtained from a cell bank of aplurality of donors. Preferably, each donor is less than 50 years of ageand/or has not been diagnosed with a vascular disease. The cells can beisolated from human aorta. Preferably, the cells can be isolated fromhuman thoracic aorta. More preferably, the cells comprise smooth musclecells.

Preferably, the cells are at passage 5 or less, at passage 4 or less, atpassage 3 or less. The cells can be cultured for a culture period ofabout six to about 11 weeks. The cells can be cultured in mediumcomprising high glucose, insulin, bFGF and/or EGF. Preferably, themedium comprises DMEM. Preferably, the cells are at cultured in mediumcomprising about 11% to about 30% human serum for the first 2-6 weeks ofculture and in medium comprising about 1% to about 10% human serum forthe remainder of the culture period (at least 4 weeks, at least 5weeks). More preferably, the cells are at cultured in a bioreactor.

The cells can be at seeded onto the tubular biodegradable polyglycolicacid construct at about 0.5×10⁶ cells per cm length of construct toabout 2×10⁶ cells per cm length of construct. Preferably, each cell ofthe seeded cells, or each cell's collective progeny, produces greaterthan 1 ng of hydroxyproline over 9 weeks in culture.

Unless otherwise defined, all technical and scientific terms used hereinhave the same meaning as commonly understood by one of ordinary skill inthe art to which this invention belongs. In the specification, thesingular forms also include the plural unless the context clearlydictates otherwise. Although methods and materials similar or equivalentto those described herein can be used in the practice or testing of thepresent invention, suitable methods and materials are described below.All publications, patent applications, patents and other referencesmentioned herein are incorporated by reference. The references citedherein are not admitted to be prior art to the claimed invention. In thecase of conflict, the present specification, including definitions, willcontrol. In addition, the materials, methods and examples areillustrative only and are not intended to be limiting.

Other features and advantages of the invention will be apparent from thefollowing detailed description and claims.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic illustration of the approach used to producereadily available extracellular matrix protein constructs. Eachconstructs is generated in the laboratory by (step A) culturing humancells on a polymer scaffold that degrades as the cells produceextracellular matrix proteins to form (step B) a tissue. Cellularmaterial is then removed, leaving (step C) an extracellular matrixconstruct, which may be refrigerated or stored at room temperature, orby some other storage means until the time of patient need. Cell-derivedextracellular matrix protein constructs may be implanted without cells(step D, diameters ≧6 mm), or (step E) seeded with recipient endothelialcells for small diameter (3-4 mm) applications.

FIGS. 2A-2G are photographs showing implant sites and observations. FIG.2A shows a human cell-derived 6 mm construct (g) implanted between theaxillary artery (a) and the brachial vein (v) in a baboon model. FIG. 2Bshows an arteriovenous graft (g) first accessed with 16 G needles at 4weeks post implant. FIG. 2C shows a representative explant angiogramdemonstrating a patent graft (g). The arterial anastomosis (aa), venousanastomosis (va), and brachial vein (v) are denoted. FIG. 2E shows acanine cell-derived 3 mm construct (g) as a carotid bypass, with clipsoccluding the intervening carotid artery (ca), at implant. FIG. 2E showsa representative angiogram demonstrating patency with no luminalnarrowing at one year. FIG. 2F shows a canine cell-derived 3 mm-diameterconstruct (g) implanted on the heart. FIG. 2G shows a CT scandemonstrating a patent graft (g) with no dilatation at 1 month.

FIGS. 3A-3F is a diagram and photographs showing decellularized humanconstructs, pre-implant. FIG. 3A shows a drawing of a 6 mm-diameterdecellularized human extracellular matrix protein construct beforeimplant. FIG. 3B shows a representative decellularized constructdemonstrating the presence of no cells in H&E stained sections (arrowpoints to residual PGA), and the porous structure typical of adecellularized construct. FIG. 3C shows a decellularized constructstained strongly and diffusely for Collagen Type I. FIG. 3D shows adecellularized construct stained for organized Collagen Type III. FIG.3E shows a decellularized construct stained for organized Fibronectin.FIG. 3F shows a decellularized construct stained for organizedVitronectin. Areas staining positive for extracellular matrix proteinsare noted with open arrowheads. In 3C-F, circumferential alignment ofextracellular matrix proteins is apparent. Note that DAB staining masksthe porous structure in Fig. C-F. Scale bars, 100 μm.

FIGS. 4A-4I are photographs showing explanted constructs remodeled invivo. FIG. 4A shows a 6 mm-diameter human extracellular matrix proteinconstruct explanted from the baboon model at 6 months demonstratingformation of a loose external adventitial-like layer (g: graft, a:“adventitia”). FIG. 4B shows a 4 mm-diameter canine construct explantedfrom carotid bypass model at 1 year (arrow points to anastomotic sutureline). FIG. 4C is a Movat's stain illustrating elastin (black) in a6-month baboon explant. FIG. 4D is an H&E stain of a 6 mm-diameterbaboon explant at 6 months showing cells densely populating graft wallsclose to the arterial anastomosis (arrowheads point to stained cells inD-I). FIG. 4E shows that after 6 months in the baboon model, α-smoothmuscle actin positive cells (brown) populated the construct wall nearanastomotic sites (note: concentrated staining was observed below theluminal surface, but cells on the lumen were not actin-positive). FIG.4F shows that these cells started to infiltrate the construct midgraftfrom surrounding adventitial-like tissue (arrows define depths of graftwalls in FIG. 4F-4H). FIG. 4G shows that in the canine model, α-smoothmuscle actin positive cells (green) were observed infiltrating intomidgraft sections of canine carotid artery bypass grafts fromsurrounding adventitial-like tissue at 1 month. FIG. 4H shows that atone year, actin-positive cells were observed through the depth of caninegraft walls. FIG. 4I shows a construct explanted from a baboondemonstrating positive staining for von Willebrand Factor in luminalcells (section near anastomosis shown). Scale bars, 100 μm.

FIGS. 5A-5F are photographs and a graph showing that the extracellularmatrix protein constructs were not immunogenic. FIG. 5A shows thatintradermal injections of homogenized graft material and PBS (negativecontrol) in baboons at 4 weeks post implantation displayed no visibleinduration or redness. FIG. 5B is a graph illustrating representativeproliferation of T-cells isolated at implant (week 0) and 24 weeks afterimplant, after exposure to segments of PTFE (negative control; notimplanted) and constructs (labeled TEVG), demonstrated that grafts areimmunologically tolerated. FIG. 5C is a photograph of H&E stainingshowing a large population of infiltrated cells in anastomotic sectionsof constructs at 6 months in the baboon. FIG. 5D shows that only asparse population of cells in anastomotic sections stain positive forCD3 (T-lymphocyte marker) at 6 months in the baboon. FIG. 5E shows thatonly a sparse population of cells in anastomotic sections stain positivefor CD20 (B-lymphocyte marker) at 6 months in the baboon. FIG. 5F showsthe absence of calcification (lack of red color) in alizarin red stainin a human constructs explanted from the baboon model at 6 months.Arrows point to stained cells. Scale bars, 300 μm.

FIG. 6 is a schematic illustration in which one donor's cells are usedto produce many extracellular matrix protein constructs for manyrecipients.

FIGS. 7A-7D is a series of photographs and diagrams showing a tubularpolymeric construct. FIG. 7A illustrates uniform PGA density. FIG. 7Billustrates non-uniform PGA density, with low and high density regions.FIG. 7C illustrates a tubular polymeric construct with a uniformlyentangled seam which matches the overall density of the tubularconstruct and a PET anchor. FIG. 7D illustrates a tubular polymericconstruct where the seam is poorly entangled and having variable density(high and low density regions).

FIGS. 8A-B are photographs of venous intimal hyperplasia. FIG. 8Aillustrates venous anastomosis of an extracellular matrix proteinconstruct of the invention. FIG. 8B illustrates substantial venousintimal hyperplasia adjacent to a PTFE graft.

FIGS. 9A-E is a series of photographs and diagrams showing usage of theextracellular matrix protein constructs of the present invention asurinary conduits. FIG. 9A shows that the conduit supports end-to-end andend-to-side anastomoses with ureters, and is tunneled in theretroperitoneal plane. At the skin, the conduit forms a stoma with theskin. FIG. 9B shows the extracellular matrix protein conduit skin stoma,with diverted urine. A stent, which is routinely used in the clinic tomaintain an open conduit during the surgical healing process, is showninserted into the extracellular matrix protein conduit through the skinstoma. FIG. 9C shows an ostomy bag collecting urine that is drainingfrom the stoma of the urinary diversion conduit. FIG. 9D shows theconduit after 28 days of exposure to concentrated urine. FIG. 9E showsthat after 4 weeks of exposure to concentrated urine, conduits resistedcrystallization, and remained physically and mechanically intact.

FIG. 10 is a diagram showing usage of the extracellular matrix proteinconstructs of the present invention as a fallopian tube conduit.

DETAILED DESCRIPTION OF THE INVENTION

The present invention provides a construct including a biodegradablepolymeric scaffold, wherein the density of the polymeric material isabout 45 mg/cc to about 75 mg/cc and said density is uniform across theentire tubular scaffold. Uniform as used herein is defined as no morethan 30%, no more than 15%, no more than 10%, no more than 5%, no morethan 4%, no more than 3%, no more than 2%, no more than 1% variabilityin density over 100% of the surface area of the scaffold. The scaffoldmay be in any shape known in the art. Preferably, the scaffold istubular. Any synthetic, biodegradable, polymeric material known in theart may be utilized. Preferably, the polymeric material is polyglycolicacid. The scaffold may be in any form known in the art. Preferably, thescaffold is felt. These constructs are referred to interchangeablyherein as “polymeric constructs”, “polymeric scaffolds”, “polyglycolicacid (PGA) constructs” or “polyglycolic acid (PGA) scaffolds”

The length of the tubular biodegradable polyglycolic acid scaffold canbe about 1 cm to about 100 cm. Preferably, the length of the tubularbiodegradable polyglycolic acid scaffold can be about 10 cm to about 40cm. More preferably, the length can be at least at least 5, at least 10,at least 12, at least 13, at least 14, at least 20, at least 25, or atleast 30 cm in length. The inner diameter of the tubular biodegradablepolyglycolic acid scaffold can be equal to or greater than about 3 mm.Preferably, the inner diameter of the tubular biodegradable polyglycolicacid scaffold can be about 3 mm to about 20 mm, such at least 3 mm, atleast 4 mm, at least 5 mm, or any integer up to about 20 mm.

The thickness of the polyglycolic acid can be about 0.8 to about 1.5 mmand said thickness is uniform across the tubular scaffold. Preferably,the polyglycolic acid can be about 0.8 to about 1.2 mm. The thickness ofthe fibers within the polyglycolic acid can be about 5 to about 20 μm.The porosity of the polyglycolic acid can be about 90% to about 98%.

The constructs of the present invention can further includenon-biodegradable polyethylene terephthalate supports at each end of thetubular biodegradable polyglycolic acid scaffold. The non-biodegradablepolyethylene terephthalate supports can be attached to the tubularbiodegradable polyglycolic acid scaffold by any means known in the art.Preferably, the polyethylene terephthalate supports are attached viasutures. The porosity of the polyethylene terephthalate can be ≧200cc/min/cm². The tubular biodegradable polyglycolic acid scaffold and thenon-biodegradable polyethylene terephthalate supports can permit theattachment and growth of cells. Alternatively, other non-degradablepolymers can be used as supports at each end of the tubular scaffold.

The constructs of the present invention are substantially free of heavymetal contaminants. Preferably, the construct includes only traceamounts of heavy metal contaminants selected from the group consistingof: aluminum, barium, calcium, iodine, lanthanum, magnesium, nickel,potassium and zinc. Aluminium can be present in an amount from about 1.5ppm to about 5 ppm. Barium can be present in an amount from about 0.03ppm to about 0.06 ppm. Calcium can be present in an amount from about 10ppm to about 4 ppm. Iodine can be present in an amount from about 0.1ppm to about 0.04 ppm. Lanthanum can be present in an amount from about0.05 ppm to about 0.3 ppm. Magnesium can be present in an amount fromabout 0.5 ppm to about 3.5 ppm. Nickel can be present in an amount fromabout 0.1 ppm to about 1 ppm. Potassium can be present in an amount fromabout 5 ppm to about 40 ppm. Zinc can be present in an amount from about1 ppm to about 5 ppm.

The constructs of the present invention can further includeextracellular matrix proteins within, and around, the biodegradablepolyglycolic acid scaffold. Preferably, the thickness of theextracellular matrix proteins is greater than about 200 micrometers atthe thinnest portion of the construct.

The present invention also provides methods of producing a tubularpolyglycolic acid construct including (a) providing a biodegradablepolyglycolic acid sheet, wherein the density of the polyglycolic acid isabout 45 mg/cc to about 75 mg/cc and the thickness of the polyglycolicacid sheet is about 0.8 to about 1.2 mm, (b) wrapping the polyglycolicacid sheet around a mandrel such that opposite edges of the polyglycolicacid sheet meet at an interface; (c) pulling polyglycolic acid fibersfrom each opposing edge of the sheet across the interface, and (d)forming a seam by entangling said pulled polyglycolic acid fibers fromone side of the interface with the polyglycolic acid fibers on theopposite side of the interface, wherein the density of the polyglycolicacid at the seam is about 45 mg/cc to about 75 mg/cc and the thicknessof the polyglycolic acid at the seam is about 0.8 to about 1.5 mm,thereby producing a tubular biodegradable polyglycolic acid constructwith a uniform polyglycolic acid density. The present invention alsoprovides a tubular biodegradable polyglycolic acid construct formed bythe methods described herein.

The mandrel can comprise any material known in the art. Preferably, themandrel comprises a gas permeable, silicone tube.

The entangling step may be performed by any method known in the artwhich permits the seam to remain intact in subsequent treatment steps.Preferably, entangling is performed with a felting needle.

The methods of the present invention can further include, treating thetubular construct to remove heavy metal contaminants. Preferably, thetubular construct is treated with one or more non-polar solvents andtreated with at least one primary alcohol, such as ethanol. Preferably,the seam remains intact following said treatment. This treatment mayalso be performed on the biodegradable scaffold prior to formation of atube.

The methods of the present invention can further include, treating thetubular construct to increase the rate of polyglycolic acid degradationand/or increase the wettability of the polyglycolic acid. Preferably,the tubular construct is treated with a strong base. More preferably,the strong base is 1M NaOH. Preferably, the seam remains intactfollowing said treatment. This treatment may also be performed on thebiodegradable scaffold prior to formation of a tube.

The methods of the present invention can further include, attachingnon-biodegradable polyethylene terephthalate supports at end of thetubular biodegradable polyglycolic acid scaffold. These supports may beattached prior to or after other treatments of the tubular construct.

The present invention also provides a tubular construct comprisingextracellular matrix proteins and a polymeric material having aninternal diameter of ≧3 mm, wherein the construct is immune andcalcification resistant, wherein the polymeric material comprises lessthan 33% of the cross-sectional area of said construct and wherein theconstruct is substantially acellular comprising less than 5% cells.Stimulation of immunity is determined, in some embodiments, by reactionto intradermal injections of construct material into the recipient, at48 hours after injection. Preferably, the polymeric material ispolyglycolic acid. These constructs are referred to interchangeablyherein as “extracellular matrix protein constructs”, “decellularizedconstructs”; or may be referred to as “grafts”, “conduits” or “vessels”depending upon in vivo usage.

The extracellular matrix protein constructs can be used in a number ofanatomical locations and disease situations. The construct can beselected from the group consisting of an arteriovenous graft, a coronarygraft, peripheral artery bypass conduit, fallopian tube replacement, anda urinary conduit. For example, extracellular matrix protein constructsare useful as arteriovenous grafts in patients undergoing hemodialysis;as coronary grafts in bypassing a blockage in patients, to bypass adiseased peripheral artery in a patient with peripheral artery disease(PAD) or as a urinary conduit. The diameter and length of theextracellular matrix protein constructs will vary for these differentuses as will the surgical attachment points. For example, a coronarygraft will attach to coronary artery, a peripheral artery graft willattach to a peripheral artery and a urinary conduit will typicallyconnect the ureter(s) to the skin to form a stoma.

Every year in the US, approximately 10,000 patients undergo cystectomy,and require a urinary conduit to drain urine outside the body(Healthcare Cost and Utilization Project (2007). N.I.S.). In almost allcases, bowel is harvested from the patient to form either an incontinenturinary diversion, or a continent urinary diversion that is catheterizedintermittently to drain urine through a continent stoma (Konety B R,Joyce G F, Wise M (2007) Bladder and upper tract urothelial cancer. JUrol 177:1636-1645). Patients may suffer from complications at the bowelharvest site, including anastomotic leaks and peritonitis. In addition,ileal urinary conduits may suffer from ischemia and necrosis, which canlead to perforation, anastomotic breakdown, stoma problems, and leakageof urine from the conduit. In the long term, many patients suffer fromchronic hyperchloremic metabolic acidosis, due to resorption of urineelectrolytes through the conduit wall. Since ileal conduits harborbacteria, patients also commonly suffer from recurrent urinary tractinfections and pyelonephritis, as bacteria from the conduit infect themore proximal urinary system. Hence, there is a significant medical needfor an improved method for urinary diversion that avoids many of thecomplications associated with the use of ileal conduits (Konety B R,Allareddy V (2007) Influence of post-cystectomy complications on costand subsequent outcome. J Urol 177:280-287; Dahl D M, McDougan W S(2009) Use of intestinal segments and urinary diversion. In: Wein A J,Kavoussi L R, Novick A C (eds) Campbell-Walsh Urology. 9th Edn.).

Surprisingly, the extracellular matrix protein constructs of the presentinvention provide significant superior properties when compared toautologous ileum. For example, no resection of the patient's intestineis needed, as surgery on the bowel is completely avoided. As describedherein, the urinary conduit of the present invention is pre-manufacturedand stored, making it readily available to patients. Since theextracellular matrix protein construct of the present invention whenused as a urinary conduit does not actively absorb its luminal contents,the risk of hyperchloremic metabolic acidosis is substantially reduced.Since the extracellular matrix protein construct of the presentinvention when used as a urinary conduit does not harbor intestinalflora, the risks of recurrent urinary tract infections are markedlyreduced. The extracellular matrix protein construct of the presentinvention when used as a urinary conduit does not produce mucus, andtherefore, risk of clogging the stoma is reduced when compared to themucus-producing ileal conduit. Since the extracellular matrix proteinconstruct of the present invention is non-living, there is essentiallyno risk of tissue ischemia due to inadequate vasculature. Rather, hostcells gradually migrate into the acellular conduit and concurrently forma microvascular network. Without ischemia, the risk of stomal stenosisis reduced. Since the extracellular matrix protein construct of thepresent invention can be grown at diameters ranging from 3-20 mm orgreater, and with lengths of up to 100 cm, it is possible to produceurinary conduits having the dimensions most suitable for diversion ofurine to the skin surface. The urinary conduit tolerates chronicexposure to urine and resists active diffusion of urine through theconduit wall.

Fallopian tube scarring is a major problem that can cause infertility.Infection, such as that associated with some sexually transmitteddiseases, can cause scar tissue to form in the fallopian tubes. Scartissue, in turn, blocks or damages the fallopian tube. A blockedfallopian tube prevents fertilization of the egg, and a damagedfallopian tube can lead to ectopic pregnancies. In the United States,more than 750,000 women experience an episode of acute pelvicinflammatory disease each year, and 10-15% of these women becomeinfertile as a result (Pelvic Inflammatory Disease (PID)—CDC Fact Sheet,National Center for HIV/AIDS, Viral Hepatitis, STD, and TB Prevention,Division of STD Prevention, September 2011). Fallopian tube canceraffects approximately 550 women in the United States each year(Vapiwala, N, and Hill-Kayser, C, Fallopian Tube Cancer: The Basics,OncoLink, Abramson Cancer Center of the University of Pennsylvania,2010). FIG. 10 shows, in a porcine model, that the extracellular matrixprotein constructs of the present invention can be sewn an end-to-endanastomosis to replace an excised segment of fallopian tube.

The tubular construct is decellularized such that it is substantiallyacellular such that the construct is immune-resistant and/orcalcification resistant. Preferably, the construct is substantiallyacellular comprising less than 2% cells, less than 1% cells or containsno cells. The cells are intact cells. The cells can be living cells ordead cells.

The tubular construct is treated to minimize the amount of polymericmaterial present. Preferably, the polymeric material is polyglycolicacid. The polymeric material may degrade or be removed such that lessthan 50%, less than 45%, less than 40%, less than 35%, less than 33%,less than 30%, less than 25%, less than 20%, less than 15%, less than10%, less than 5%, less than 3%, or less than 1% of the cross-sectionalarea of the tissue comprises the synthetic polymeric material.Calculation of the cross-sectional area does not include the lumen.

The length and diameter of the extracellular matrix protein constructmay vary with the anatomical application desired. The length of theextracellular matrix protein construct acid scaffold can be about 1 cmto about 100 cm. Preferably, the length is about 10 cm to about 40 cm.More preferably, the length can be at least at least 5, at least 10, atleast 12, at least 13, at least 14, at least 20, at least 25, or atleast 30 cm in length. The inner diameter of extracellular matrixprotein construct can be equal to or greater than about 3 mm.Preferably, the inner diameter is be about 3 mm to about 8 mm, such atleast 3, at least 4, at least 5, at least 6, at least 7, or at least 8mm. More preferably, the inner diameter can be about 3 mm to about 20mm.

Surprisingly, the extracellular matrix protein construct can comprise aburst pressure of at least 600 mm, at least 700 mm, at least 800 mm, atleast 900 mm, at least 1000, at least 1100 mm, at least 1200 mm, atleast 1300 mm, at least 1400 mm, at least 1500 mm, at least 1600 mm, atleast 1700, at least 1800 mm, at least 1900 mm or at least 2000 mm Hg.Burst pressure can be measured by any means known in the art; forexample, by inflating the construct with a fluid at gradually increasingpressures, until the construct either ruptures or forms a discrete hold.Preferably, the construct has a burst strength greater than 2000 mm Hg.Equally surprising, the construct can comprise a suture strength ofgreater than 60 g, 70 g, 80 g, 90 g or 120 g. Preferably, the constructhas a suture strength greater than 120 g. Suture strength can bemeasured by any means known in the art; for example, by inserting a 6-0suture through the construct at a distance of 2 mm from the edge of theconstruct. The thickness of the tubular construct can be greater thanabout 200 micrometers at the thinnest portion of the construct. Theconstruct can be impermeable to fluid. The fluid can be saline or abiological fluid such as blood or urine. Impermeable to fluid is definedas the absence of net fluid flow out of the construct after filling withfluid at atmospheric pressure, at 200 mm Hg, at 300 mm Hg or 400 mm Hg.Preferably, the construct is impermeable to fluid leakage up to at least200 mm Hg, at least 300 mm Hg or at least 400 mm Hg. Given that normalpressures in blood vessels do not exceed 120 mm Hg, and that severeStage 4 blood pressures in hypertension reach a maximum of up to 230 mmHg, the extracellular matrix protein constructs provided herein resistleakage and weeping in both healthy and sick patients. Given thaturetheral pressures are approximately 30 mmHg, this construct alsoresists leakage of urine during use as a urinary conduit. Without beingbound by any theories, the extracellular matrix protein constructsprovided herein resist leakage and weeping because the extracellularmatrix proteins (e.g., collagen) is closely packed in the constructs,with a density of 856±221 micrograms hydroxyproline per cm² of constructmaterial.

The extracellular matrix proteins can comprise hydroxyproline,vitronectin, fibronectin and collagen type I, collagen type III,collagen type IV, collagen VI, collagen XI, collagen XII, fibrillin I,tenascin, decorin, byglycan, versican and asporin. Preferably, theextracellular matrix proteins can comprise hydroxyproline at >40 μg/mgdry weight. Preferably, the extracellular matrix proteins are producedfrom allogeneic, autologous or xenogeneic cells. Preferably, theextracellular matrix proteins are, in part, oriented circumferentiallyaround the tubular construct. Circumferential orientation ofextracellular matrix proteins provides an “anchor” for sutures. Thearrows in FIG. 3 highlight the circumferential orientation of fibers. Incontrast, having a predominantly axial orientation of extracellularmatrix does not provide the sutures with a structure to anchor onto;rather the suture would slip through axially aligned fibers.

The construct can comprise less than 300 ng/cm of beta-actin.Preferably, the construct comprises <150 ng/cm of beta-actin. The amountof beta-actin can be determined by any means known in the art; forexample, by ELISA assay. The construct can comprise less than 3% dryweight of lipids. The amount of lipid can be determined by any meansknown in the art; for example, by gas chromatography-mass spectrometry.The construct can comprise trace amounts or no detectable amounts ofdouble stranded genomic DNA. Preferably, the amount of DNA is asdetermined by gel electrophoresis.

The construct induces little to no calcification upon implantation invivo. Preferably, the construct induces less than 1% calcificationwithin 6 weeks, within 3 months, within 6 months, within 9 months, or 12months of implantation. More preferably, the construct induces less than0.5% calcification within 6 weeks, within 3 months, within 6 months,within 9 months, or 12 months of implantation. Most preferably, theconstruct produces no calcification within 6 weeks, within 3 months,within 6 months, within 9 months, or 12 months of implantation.Calcification can be determined by any means known in the art; forexample, calcification is measured by percent area of the construct thaton histologic sectioning stains positive for calcium, using ahistochemical stain such as alizarin red stain.

The construct is immune-resistant such that the construct induces littleto no immune response upon implantation in vivo, as defined by whealformation and redness at 48 hours after intracutaneous injection ofgraft material particles into the recipient. Preferably, the constructinduces less than 1 mm of intimal hyperplasia thickening in nativevasculature at anastomoses with the construct at 3 months, 6 months, 9months, or 12 months of implantation, when implanted as a vasculargraft. More preferably, the construct induces less than 1, 0.75, 0.5,0.4, 0.3, or 0.25 mm of intimal hyperplasia thickening in nativevasculature at anastomoses with the construct at 6 months ofimplantation.

Surprisingly, the extracellular matrix protein constructs do not dilategreater than 50, %, greater than 40%, greater than 30%, or greater than20% beyond their diameter at time of implant. This is very beneficial invivo. Also surprisingly, the extracellular matrix protein constructs arevery storage stable, and can be stored before or afterdecellularization. The construct may be stored at about 2° to about 30°C. for at least 1, 2, 3, 4, 6, 8, 12, 18, or 24 months withoutcomprising their integrity and implantability. The integrity of theconstructs can be assessed by any means know in the art; for example, asassessed by retained suture retention strength of at least 80% ofstarting value. The constructs can be stored in any suitablephysiological buffer known in the art. The buffer may include proteaseinhibitors, or ion chelators. This is very beneficial as theextracellular matrix protein constructs are readily available (minimalor no wait time) for in vivo use.

The present invention also provides methods of producing a tubularconstruct comprising (a) providing a tubular biodegradable polyglycolicacid construct, (b) seeding human cells at passage 6 or less on thetubular biodegradable polyglycolic acid construct, (c) culturing thecells under conditions such that the cells secrete extracellular matrixproteins on the tubular biodegradable polyglycolic acid construct, (d)decellularizing the construct in step (c) such that the construct issubstantially acellular comprising less than 5% cells and wherein theconstruct is immune and calcification resistant, and (e) degrading thepolyglycolic acid construct in step (c) such that the polyglycolic acidcomprises less than 33% of the cross-sectional area of said construct,thereby producing a decellularized tubular construct. The presentinvention also provides a decellularized tubular construct formed by themethods described herein.

The tubular construct is decellularized such that it is substantiallyacellular such that the construct is immune-resistant and/orcalcification resistant. Preferably, the construct is substantiallyacellular comprising less than 2% cells. More preferably, the constructis substantially acellular comprising less than 1% cells. Mostpreferably, the construct contains no cells. Thus, in thedecellularization step greater than 25%, greater than 40%, greater than50%, greater than 75%, greater than 85%, greater than 90%, greater than95%, greater than 98%, or greater than 99% of the cells seeded onto thetubular biodegradable polyglycolic acid construct are removed. The cellscan be allogeneic, autologous or xenogeneic to the host to which theconstruct will be implanted. Preferably the cells are allogeneic.

The cells are obtained from a single donor or obtained from a cell bank,wherein the cells in the cell bank are pooled from a plurality ofdonors. Preferably, the cells are obtained from a cell bank of aplurality of donors. Preferably, each donor is less than 50 years of ageand/or has not been diagnosed with a vascular disease. The cells can beisolated from human aorta, femoral artery, iliac artery, carotid artery,radial artery, ureter, bladder wall or skin. Preferably, the cells areisolated from human aorta. More preferably, the cells are isolated fromhuman thoracic aorta. The cells can comprise smooth muscle cells,mesenchymal cells, fibroblasts, fibrocytes, and/or endothelial cells.Preferably, the cells comprise smooth muscle cells.

The seeded cells are low passage cells, having been passaged less than10, less than 5, less than 5, less than 4, less than 3, or less than 2times. Preferably, the cells are at passage 3 or less. The cells can becultured for a culture period of about six to about 11 weeks. Thepolymeric scaffold may be in any form during the phase of culturing. Itcan be in the ultimate shape, or it can be shaped after the phase ofculturing. Preferably, the scaffold is tubular. Alternatively, thescaffold is shaped into a tubular form after the culturing. Culturing ofthe cells can be performed using any conventional medium and apparatus,taking into account nutritional, oxygenation, temperature, mechanical,and pressure conditions. The medium may optionally comprise bovineserum, porcine serum, ovine serum, equine serum, or human serum. Suchsera may provide growth factors and known or unknown components forimproving properties of the culturing process. As the cells grow inculture on the scaffold, they secrete collagenous extracellular matrix.Preferably, the cells are at cultured in medium comprising 20% humanserum for the first 2-6 weeks of culture and 10% human serum for theremainder of the culture period. More preferably, the cells are atcultured in a bioreactor.

The cells can be at seeded onto the tubular biodegradable polyglycolicacid construct at about 0.5×10⁶ cells per cm length of construct toabout 2×10⁶ cells per cm length of construct. Preferably, each cell ofthe seeded cells, or each cell's collective progeny, produces greaterthan 1 ng of hydroxyproline over 9 weeks in culture.

The methods of the present invention can further include: prior toimplantation in a patient, seeding cells on the decellularized tubularconstruct. The cells can include smooth muscle cells or endothelialcells. Preferably, the cells are endothelial cells. The cells can beallogeneic or autologous cells. Preferably, the cells are autologouscells. Most preferably, the cells are autologous endothelial cells.

Decellularized extracellular matrix protein constructs have severaladvantages over decellularized human cadaveric vessels. First, cadaverichuman vasculature has small branches that must be ligated, whereasengineered tissues consist of a tube without branches. Second,decellularized extracellular matrix protein constructs have a loosetissue structure without layers of lamellar elastin. This loosestructure allows decellularization solutions to readily permeateengineered tissues to remove cellular material without excessiveexposure that may damage extracellular matrix integrity, and also mayimprove cellular repopulation in vivo. Thirdly, using a decellularizedextracellular matrix protein constructs approach maximizes the impact ofhealthy tissue donors by allowing production of a large number of graftsper donor, whereas a decellularized cadaveric vascular graft approachhas limited amounts of available vascular tissue per donor withdiameters that are appropriate for common cardiovascular surgicalprocedures. Extracellular matrix protein constructs can be created in avariety of diameters that can more suitably match bypassed nativearterial vasculature. In contrast, decellularized human cadavericvessels cannot be created for a particular diameter, and size mismatchbetween the small native vessel and large bypass graft can occur,potentially resulting in diminished patency rates.

Constructs can be decellularized using any means known in the art; forexample, as described previously (Dahl, et al., Cell Transplantation 12,659-666 (2003)). One preferred decellularization solution comprisesphosphate buffered saline (PBS) with 0.12M sodium hydroxide, 1M sodiumchloride, and 25 mM EDTA, containing either 8 mM CHAPS or 0.07-1.8 mMSDS. Another preferred decellularization solution does not comprise SDS.The decellularization methods of the present invention may include abenzonase step to digest DNA. Preferably, the benzonase step includes asolution comprising 2 U/mL Benzonase, 47 mM Tris, 1.4 mM magnesiumchloride, and 19 mM sodium chloride, at pH 8.0.

The presence of sparse residual PGA fragments in extracellular matrixprotein constructs at the time of implant is not of concern, as PGA isan FDA-approved degradable suture material with breakdown products thatare readily metabolized. Further, PGA has been used as a vascular graftcomponent without any known negative effects on vascular remodeling(Shin'oka, et al., J Thorac Cardiovasc Surg 129, 1330-1338 (2005)). Thehuman cell-derived grafts produced in this study were an order ofmagnitude stronger than those described in previous reports that alsoused PGA as a support for tissue creation (Poh, et al., Lancet 365,2122-2124 (2005); McKee, et al., EMBO Rep 4, 633-638 (2003)). However,it is important to note that these prior reports utilized human venouscells or commercially available human aorta cells at high passage (Poh,et al., Lancet 365, 2122-2124 (2005); McKee, et al., EMBO Rep 4, 633-638(2003)). In previous reports, use of dense PGA sutures to sew sheets ofPGA into tubes left a substantial amount of residual PGA inextracellular matrix protein constructs, which diminished extracellularmatrix protein construct strengths (Dahl, et al., Ann Biomed Eng 35,348-355 (2007)).

In large diameter applications, such as above-the-knee peripheral bypasssurgery and hemodialysis access, PTFE vascular grafts function wellenough to warrant routine clinical use (Harris, et al., J Vasc Surg 33,528-532 (2001)). Therefore, large diameter extracellular matrix proteinconstructs can be utilized without luminal EC seeding. However, forsmall diameter applications, it has been extremely difficult to find afunctional vascular graft other than the patient's own vasculature(Harris, et al., J Vasc Surg 33, 528-532 (2001)), which is highlycompliant (Table 3) and contains ECs. To minimize risk of graftocclusion, ECs were seeded onto extracellular matrix protein constructsprior to implant in the small diameter peripheral and coronary settingsto provide an antithrombogenic luminal surface. ECs were isolated fromperipheral arteries or veins of dogs prior to undergoing bypass withextracellular matrix protein constructs. This is similar to peripheralvein harvest approaches previously reported for isolation of ECs forvascular graft seeding (McAllister, et al., Lancet 373, 1440-1446(2009); Deutsch, et al., J Vasc Surg 49, 352-362 (2009)). Autologous ECscould also be isolated more rapidly from adipose tissue (Arts, et al.,Lab Invest 81, 1461-1465 (2001)) or circulating blood (Kalka, et al.,Proc Nat Acad Sci 97, 3422-3427 (2000); Hill, et al., New Eng J Med 348,593-600 (2003)), which could reduce the patient wait time forendothelialization from weeks to days or possibly even to hours.

The observed patency rate of 83% for small diameter extracellular matrixprotein constructs with poor luminal EC coverage suggests that completeluminal EC coverage prior to implant is not be required for graftfunction in the setting of systemic anti-platelet therapy throughout theduration of implantation. Poor EC coverage at implant is also observedin saphenous vein grafts, which are often denuded of endothelium duringgraft isolation (Roubos, et al., Circulation 92, 31-36 (1995)). It ispossible that the presence of sparse ECs at the time of implant aids inmaintaining patency in vivo, either by supplying sufficient release ofanti-thrombogenic signals or by aiding in recruitment of recipient ECsto the extracellular matrix protein construct luminal surface (Lee, etal., Circulation 114, 150-159 (2006)). On the other hand, extracellularmatrix protein constructs may be less thrombogenic than other, syntheticvascular graft materials and may function without ECs on the luminalsurface at the time of implant.

The functional effects of immunogenicity (intimal hyperplasia,aneurysmal dilatation, or calcification in the long term (Sclafani, etal., Arch Facial Plast Surg 2, 130-136 (2000); Mitchell and Libby, CircRes 100, 967-978 (2007); Yankah and Wottge, J Card Surg 12, 86-92(1997))) were not observed in baboon or canine studies, demonstratingthat the disclosed tissue engineered vascular grafts werenon-immunogenic. In contrast, discordant xenogenic extracellular matrixproteins and allogeneic cells (found in bovine vascular xenografts andhuman cadaveric cryopreserved vascular allografts, respectively) triggerimmunological responses and their functional side-effects (Allaire, etal., Surgery 122, 73-81 (1997); Carpenter, and Tomaszewski, J Vasc Surg27, 492-499 (1998)). Extracellular matrix protein constructs resistedintimal hyperplasia formation in long-term implants. Extracellularmatrix protein constructs demonstrated less neointimal hyperplasia at 6months as arteriovenous grafts than PTFE at 1 month as arterial bypassgrafts (Lumsden, et al., J Vasc Surg 24, 825-833 (1996)), which isencouraging given that arteriovenous grafts typically trigger moresubstantial intimal thickening than do arterial bypass grafts. Giventhat end-to-side carotid artery bypass has been described as a modelthat results in extensive intimal hyperplasia at one month (Kapadia, etal., J Surg Res 148, 230-237 (2008)), the absence of intimal hyperplasiaat one year in the canine peripheral bypass studies is surprising.

The extracellular matrix protein constructs of the present invention areavailable without a significant patient wait time and represent asubstantial advance over completely autologous tissue engineeringapproaches, wherein patients must wait for long time periods for graftsto be cultured. The constructs provided herein are functional asarteriovenous conduits, and as small-caliber arterial bypasses in theperipheral (carotid) and coronary circulations. Conduits used previouslyin the clinic have suffered from substantial intimal hyperplasia,aneurysm, and calcification. Encouragingly, the decellularizedextracellular matrix protein constructs resist substantial intimalhyperplasia, dilatation, and calcification in various large-animalmodels. These data support the use for the decellularized humanextracellular matrix protein constructs in a range of vascularapplications for patients who have no available autologous vascularconduit.

Examples are provided below to further illustrate different features ofthe present invention. The examples also illustrate useful methodologyfor practicing the invention. These examples do not limit the claimedinvention.

Example 1

Formation of Polymeric Scaffold:

Measure proper width and length of the PGA mesh (Polyglycolic acid felt)required and cut to size. For example, 3 mm-1.35 cm×desired length; 4mm-1.66 cm×desired length; or 6 mm-2.35 cm×desired length. PGA Mesh canbe obtained from Biomedical Structures (1 mm thick, 50 mg/cc (Range45-58), 20×30 cm). Wrap mesh around appropriately sized silicone tubingcut 10 cm longer than length of mesh. Use felting needle to pull a fiberthread from one side of the mesh to the other side to entangle PGAfibers along the seam. Repeat all along seam edge. Entangle fiberstightly against silicone tubing to create a vessel/tubular shape of themesh. Seam should be no thicker than the rest of the tube. The seam isthen secured by mending any tears, holes or thin spots throughout meshtube

The PGA is ideally 45-75 mg/cc. Low density (<45 mg/cc) regions lack asufficient number of PGA fibers to entangle a seam without holes. Lowdensity regions also lead to reduced cell attachment, and poor cellattachment may lead to insufficient local extracellular matrixproduction. High density (>75 mg/cc) PGA is associated with a greaterdensity of PGA residuals in the final product. FIG. 7A shows auniform-density PGA felt with density in the range of 45-75 mg/cc. FIG.7B shows a non-uniform-density PGA felt, with regions of unacceptablylow density (<45 mg/cc) that are unacceptable for use and regions ofhigh density (>75 mg/cc) that may lead to increased residual PGA in thefinal product.

The fiber entangling method is used to turn PGA sheets (see PGA sheet inFIG. 7A) into tubes. The entangling method involves wrapping a strip ofPGA around the silicone mandrel and meeting the edges of PGA at aninterface. Thereafter, fibers of PGA from each side of the strip arepulled across the interface and inserted between fibers on the oppositeside of the interface. A sufficient number of fibers must be pulled tomake the “seam” strong enough to withstand subsequent scouring andsurface treatment. The fibers must also be pulled in such a manner thatthe seam density matches that of the rest of the tubular scaffold (seeFIG. 7C), so that cells will be distributed uniformly around the PGAtube and will produce a uniform tissue. If the seam is not uniform (seeFIG. 7D), areas of very low density will become holes during the NaOHsurface treatment process. In addition, low-density areas in the PGAseam may lead to poor local cell seeding, which may lead to a thin spotin the final graft. High density areas in the seam as shown in FIG. 7Dmay locally increase PGA residuals in the final product. Locallyconcentrated residuals of PGA in grafts may locally reduce graftstrength (Dahl et al. Ann Biomed Eng 35 (3):348-355 (2007))

Cut polyethylene terephthalate (PET) material (Dacron® material) into 1cm segments—about 6-7 ribs. Dacron® material can be obtained from Maquet(Rastatt, Germany; Product No. 174408, C-Code C1768, D 8 mm×L 50 cm,average porosity 260 cc/min/cm²). Since the Dacron® does not gathereasily, cut a small triangular wedge into Dacron cuff to fit 3 or 4 mmtubing. When the Dacron® is sutured to the mesh, close the triangularwedge to fit snuggly to the mesh and silicone tubing. Cut small slice,about 3 ribs, into Dacron® cuff to accommodate 6 mm tubing attachment tobioreactor. Using 4.0 Surgipro™ II suture (Coviden Ltd, Dublin, IRL),first sew wedge shut to fit 3 or 4 mm tube, then attach cuff to PGA meshtube using surgeons' stitches. Sew running stitches across top of cuffto create purse string closure onto bioreactors. The tubular PGAconstruct may be stored or treated as described below.

The tubular configuration of PGA shown in FIG. 7C allows cells to seedand thereafter grow in a tubular shape. A polyethylene terephthalate(PET) cuff shown in FIG. 7C supports ingrowth of tissue and therebybecomes integrated with the growing tissue. PET's non-degradableproperty allows it to serve as an anchor to the bioreactor to hold thetissue at a fixed length during culture. In contrast, the PGA tubedegrades during the tissue growth phase. The inner diameter of the PGAtube is defined by the outer diameter of the mandrel around which thePGA scaffold is entangled, and in this case, the mandrel is made ofsilicone. FIG. 7C shows a silicone mandrel with outer diameter of 6 mm,and the PGA tube formed around the silicone mandrel has an innerdiameter of 6 mm. As tissue forms, the contractile cells contract thepolymer and tissue around the silicone tube such that the resultanttissue inner diameter is also defined by the outer diameter of thesilicone tube Inner diameter of PGA tube (and outer diameter of thesilicone tube mandrel) produced readily is in the range of 3-6 mm, andtissues with smaller or larger diameters may be created by usingsilicone tube with the desired diameter.

A scour process is used to remove heavy metals, lubricants, and othercontaminants. The PGA tube is placed on a mandrel and washed with moreor more non-polar solvents and at least one primary alcohol, such asethanol for at least 30 minutes while shaking at 25 rpm. Dry PGA tubesovernight.

Table 1 shows that this scour method substantially removes heavy metalcontaminants. The presence and/or amount of heavy metal contaminants canbe determined by any means know in the art; for example, by massspectroscopy.

TABLE 1 Pre-Scour, PPM Post-Scour, PPM Aluminum 15 2.3 Barium 0.79 <0.05Calcium 35 <7 Iodine 170 <0.07 Lanthanum 5 <0.1 Magnesium 4.5 <2 Nickel3.5 0.31 Potassium 120 <20 Zinc 11 2.7

Treatment of PGA with NaOH has been shown to increase the rate of PGAdegradation. For example, no NaOH treatment resulted in 50% mass losswithin 8.5 weeks, but 1-3 minutes of 1M NaOH treatment resulted in65-70% mass loss within 8.5 weeks (Prabhakar et al., (2003) Engineeringporcine arteries: effects of scaffold modification. J Biomed Mat Res67A:303-311).

The PGA tube degrades as the tissue is cultured in the bioreactor. Asthe PGA degrades, cells are producing extracellular matrix proteins thatform the bulk of the final graft mass. As an example, consider thefollowing:

PGA Time 0, prior to bioreactor culture:

-   -   PGA sheet density: 55 mg/cc    -   Time 0 volume of PGA in a 6 mm-diameter graft: 0.235 cc/cm tube    -   Time 0 density of PGA in a 6 mm-diameter graft is then: (55        mg/cc)*(0.235 cc/cm tube)=12.9 mg PGA/cm graft.

PGA in the final extracellular matrix protein construct, followingbioreactor culture and decellularization as described below:

-   -   PGA mass loss after 8.5 weeks in an aqueous environment at 37°        C.: 30% PGA mass remains after 8.5 weeks.    -   Mass PGA per cm graft: (30%)×(12.9 mg PGA/cm graft)=3.9 mg        PGA/cm graft        -   Average wet weight of a 6 mm graft: 600 mg graft/cm graft        -   Mass PGA per mass graft: (3.9 mg PGA/cm graft)/(600 mg            graft/cm)=0.0065 mg PGA/mg graft.

Therefore, PGA constitutes <1% of the final hydrated mass in thisexample. Similarly, using the maximum PGA density specification of 75 mgPGA/cc PGA in the same set of calculations would produce a graft withPGA as <1% of final graft mass (0.0088 mg PGA/mg graft).

Example 2

Animal Use:

All procedures were approved by their respective Animal Care and UseCommittees, including Duke University, East Carolina University, andSyneCor. Animals received humane care according to the “Guide for theCare and Use of Laboratory Animals” (NIH, 1996). All surgeries andangiography were performed in sterile fashion under general anesthesia.After each surgery, graft patency was confirmed, wounds were closed, andanimals were recovered. Animals were anti-coagulated with heparin(1000-5000 U) at implant. Baboons received aspirin (10 mg/kg), and dogsreceived dual anti-platelet therapy (325 mg aspirin/75 mg clopidogrel),daily preoperatively until the end of the study.

Formation of Extracellular Matrix Protein Constructs:

Human aortas were obtained from an American Association of Tissue Banks(AATB) accredited and FDA registered tissue bank (CryoLife, Inc.;Kennesaw, Ga.), and met criteria for implantation (FDA 21CFR1271, AATBStandards for Tissue Banking, and internal CryoLife acceptancecriteria). Human smooth muscle cells (SMCs) were isolated from donoraortas (ages 17-49) that were consented for research use and tested forbioburden (aerobic bacterial and fungal contaminants), sterility,mycoplasma, and endotoxin. Cells were stored in liquid nitrogen vapor(−135° C.) prior to use. Cells from multiple donors were pooled forculture of pooled donor grafts. Human cells were used at passage 2.

Canine SMCs were isolated from canine carotid and femoral arteries, wereallogeneic with respect to the construct recipient, and were used atpassage 2-4.

Using an aseptic process, cells (either human or canine) were seededonto tubular poly-glycolic acid felt scaffolds (6 mm ID for humanconstructs, and 3 or 4 mm ID for canine constructs) and strainedcyclically (2.5% at 2.75 Hz) (Niklason, et al., Science 284, 489-493(1999)) in a bioreactor to produce construct. Medium for growth of humanconstructs was high glucose DMEM with 20% serum, 5 milligram insulin perL, 5 microgram bFGF per L, 1 microgram EGF per L, 100,000 U penicillin Gper L, 3 micrograms copper sulfate per L, 50 milligrams L-proline per L,40 milligrams L-alanine per L, 50 milligrams glycine per L, and 50milligrams ascorbic acid per L, and was changed thrice weekly. Mediumfor growth of canine constructs was low glucose DMEM with 20% serum, 10ng/ml PDGF-BB, 10 ng/ml bFGF, 500 U/ml penicillin G, 3 ng/ml coppersulfate, 50 ng/ml L-proline, 20 ng/ml L-alanine, and 50 ng/ml glycine,and was changed once per week. L-ascorbic acid was added thrice weeklyto canine extracellular matrix protein construct cultures.

After 7-10 weeks of culture, constructs were decellularized usingaseptic processing. As described previously (Dahl, et al., CellTransplantation 12, 659-666 (2003)), decellularization solutionscomprised phosphate buffered saline (PBS) with 0.12M sodium hydroxide,1M sodium chloride, and 25 mM EDTA, containing either 8 mM CHAPS or0.07-1.8 mM SDS. An alternative decellularization method was alsoemployed, in which the SDS was removed. In addition, a benzonase stepmay be added to digest DNA, using 2 U/mL Benzonase, 47 mM Tris, 1.4 mMmagnesium chloride, and 19 mM sodium chloride, at pH 8.0. Extracellularmatrix protein constructs were exposed to each solution for up to 6hours at room temperature, and were then washed with PBS. Allextracellular matrix protein constructs were decellularized prior tomechanical testing, endothelial cell seeding, and implantation.Decellularized extracellular matrix protein constructs were stored at 4°C. in phosphate buffered saline (PBS) without calcium or magnesium.

Endothelialization of Extracellular Matrix Protein Constructs:

Canine extracellular matrix protein constructs were seeded withautologous endothelial cells (ECs) in vitro prior to implantation.Canine femoral artery, carotid artery, or cephalic vein segments (3-4cm) were cultured on fibronectin-coated plates in low glucose DMEM with10% FBS, 1× microvascular growth supplement, 125 μg/ml heparin, and 500U/ml penicillin G, for isolation of ECs via outgrowth from each segment.Isolation and expansion of ECs required 21±2 days, and the attachment ofECs to extracellular matrix protein constructs and shear preconditioningrequired an additional 2 days. For EC attachment, graft lumens werecoated with fibronectin (100 μg/ml), seeded with ECs (750,000/ml), andexposed to 11 hours of rotation at 10 rotations per hour to encourageeven distribution of ECs. Shear preconditioning was performed byincreasing the mean velocity of perfused culture medium in stepwisemanner (10 steps total) over a 22-hour period, and maintaining themaximum mean velocity (10-15 cm/s) to match the mean velocity reportedfor peripheral canine arteries (10-16 cm/s (Pedley, The Fluid Mechanicsof Large Blood Vessels (Cambridge University Press, Cambridge, UK,1980)) for 13 hours prior to implant.

In Vitro Analysis:

Suture retention strengths were measured by passing a loop of 6-0Prolene® suture (BV-1 needle) through each extracellular matrix proteinconstruct, 2 mm from the edge, and suspending weights in 10 g incrementson the suture loop until the suture pulled through the tissue. Suturestrength was defined as the weight in grams required to tear the tissue.Suture strengths of extracellular matrix protein constructs weremeasured prior to implant, and after baboon explant. Burst pressureswere measured prior to implant by inflating 6 mm human extracellularmatrix protein constructs, or 3-4 mm canine extracellular matrix proteinconstructs, with saline at room temperature until rupture, as previouslydescribed (Dahl, et al., Cell Transplantation 12, 659-666 (2003)). Burstpressure was defined as the inflation pressure at which an extracellularmatrix protein construct ruptured.

For DNA quantification, extracellular matrix protein construct segmentswere digested with papain followed by DNA purification using a modifiedQiagen silica based spin column. Resulting captured DNA was eluted usinga detection compatible buffer, and DNA was quantified using a PicoGreen®assay (Life Technologies, Inc, Grand Island, N.Y., USA). Hydroxyprolinewas measured in papain-digested samples, using chloramine T andp-dimethylaminobenzaldehyde, and collagen was calculated as 10 times theamount of hydroxyproline

Animal Models:

An old world primate model was chosen to provide phylogenetic similarityto humans, which allowed implantation of non-crosslinked humanmatrix-containing grafts without immunosuppression. Adult male baboons(Papio Anubis, 20-30 kg) are physically large enough to supportimplantation of a 6 mm-diameter extracellular matrix protein constructin a clinically relevant anatomic setting. Primates, however, aresignificantly more expensive than other animals (Rashid, et al.,Biomaterials 25, 1627-1637 (2004)), are difficult to handle and maintain(Narayanaswamy, et al., J Vasc Intervent Radiol 11, 5-17 (2000)), andare limited in availability. Thus, baboons were used for arteriovenousstudies, while dogs were utilized for small-diameter investigations.

The canine model (Class A Mongrel dogs, ˜25 kg) was employed for theassessment of 3-4 mm diameter extracellular matrix protein constructsdue to its wide acceptance in the scientific community for theevaluation of vascular prostheses (Tomizawa, et al., Circulation 90(part 2), II-160-166 (1994); Bianco, et al., Large Animal Models inCardiac and Vascular Biomaterials Research and Testing. B. D. Ratner, F.J. Schoen, A. S. Hoffman, J. E. Lemons, Eds., “Biomaterials Science: AnIntroduction to Materials in Medicine” (Elsevier Science & TechnologyBooks, 2004)). The canine study utilized an allogeneic acellularextracellular matrix protein construct, seeded with autologous ECs,which mimics the approach proposed for eventual small-diameter clinicaluse

Surgical Implantation Techniques:

Nine adult male baboons underwent arteriovenous placement of humanextracellular matrix protein constructs (6 mm ID). One extracellularmatrix protein construct was placed in the aorto-caval position for onemonth. Eight extracellular matrix protein constructs were placed betweenthe axillary artery and distal brachial vein, which provided asuperficial site amenable for simulating hemodialysis access, for up tosix months. All anastomoses were created with a running 6-0 Prolene®suture technique.

To test long-term in vivo patency, canine extracellular matrix proteinconstructs (3-4 mm ID) seeded with autologous ECs were implantedend-to-side to the carotid artery in five dogs using 8-0 Prolene®suture. The intervening native carotid artery was occluded with surgicalclips.

Canine extracellular matrix protein constructs (3-4 mm ID) seeded withautologous ECs were implanted into the coronary circulation of threedogs. A left thoracotomy exposed the heart. Normothermic cardiopulmonarybypass was utilized and cardiac standstill was achieved with coldcardioplegia. Each extracellular matrix protein construct was sutured tothe left anterior descending coronary artery (8-0 Prolene®) and to theascending aorta (4.0 mm aortotomy, 7-0 Prolene®), with ligation of theproximal coronary artery. After coronary bypass, animals were separatedfrom cardiopulmonary bypass and recovered.

Immunological Assessments:

In the concordant xenogenic model of human cell-derived constructsimplanted into baboon, immunogenicity of human matrix-containingconstructs was assessed. Subcutaneous injections of homogenizedextracellular matrix protein constructs (0.1 ml of a 0.25 mg protein/mlphosphate buffered saline, PBS) and PBS negative control (0.1 ml) wereadministered at days 0 and 28, with visual assessments 48-72 hours aftereach injection to detect whether an in vivo adaptive immune response wasforming.

In addition, T-cell proliferation was measured at 0, 4, 12, and 24 weeksfor baboon implants. Lymphocytes were isolated with a Ficoll gradient,and cultured 7 days in 96-well plates with segments (5 mm×5 mm) ofextracellular matrix protein construct, or PTFE grafts as negativecontrols, in each well. Culture medium was RPMI 1640 with 10% fetalbovine serum. BrdU (100 μM) was added to each well 18 hours before cellharvest. Harvested cells were stained with 200 μL of diluted live/deaddye (Invitrogen™, Carlsbad, Calif.; L23102) for 30 min at roomtemperature and then with 80 μL of CD3-APC antibody (BD Pharmingin™, SanJose, Calif.; BD557597) for 50 min at room temperature, washed,permeabilized (1× Cytofix/Cytoperm™ buffer and 1× Cytoperm™ Plus; BDBiosciences, and incubated with 100 μL of DNase to partially digest DNA.Proliferating cells were stained with 50 μL of BrdU-FITC antibody (BD559619) for 20 min at room temperature and suspended in 150 μL of 0.2%BSA/DPBS for Flow Cytometry analysis (BD Accuri™ C6). For data analysis,single cells were selected and dead cells were gated out. Proliferationrate was calculated as the percentage of CD3+/BrdU+ cells in CD3+ cells.

Duplex Ultrasound:

In the baboon model, duplex ultrasound was used to monitor midgraftextracellular matrix protein construct diameter, wall thickness, andflow rate immediately after surgery, and at 2, 4, 12, and 24 weeks.

Angiography:

Angiography was used to assess graft dilatation and narrowing. Graftpatency was defined according to Fitzgibbon's classification(Fitzgibbon, et al., J Am Coll Cardiol 28, 616-626 (1996)).

All baboon grafts placed from the axillary artery to the brachial veinwere accessed directly in mid or distal graft sections (16 G needle, 5-6F catheters) at 1, 3, and 6-month time points (see Table 6) to determinethe ability of extracellular matrix protein constructs to withstandpuncture as a model for hemodialysis access.

Angiography of canine constructs was performed through a percutaneousfemoral arterial approach at 1, 4, 12, 26, and 52 weeks after implant.

Computed Tomography Angiography:

Computed tomography angiography (64-slice; General Electric, Lightspeed®VCT; Fairfield, Conn., USA) of coronary bypass grafts was performed.Intravenous β-blockers minimized cardiac motion, and iohexol (350 mgI/mL) was used as contrast. Slices (0.625 mm thick) and a soft-tissuereconstruction algorithm were used for evaluation of the internaldiameter and cross-sectional area of grafts.

Histology:

Tissues were fixed in 10% neutral buffered formalin, embedded inparaffin, sliced (5 μm sections), and stained with H&E, Movat's, orAlizarin Red S with a light green counterstain. Tissue sections werealso prepared for cryosectioning by dehydrating (30% sucrose inphosphate buffered saline, PBS) and freezing in Tissue-Tek® optimalcutting temperature (OCT) compound (Sakura Finetek USA, Torrance,Calif.). Immunostaining was performed on frozen baboon sections andformalin-fixed canine sections for α-smooth muscle actin (SMC andmyofibroblast marker; baboon explants: Dako M0851, 1:50 dilution (Dako,Carpinteria, Calif., USA); canine explants: Sigma A2547, 1:5000 dilution(Sigma-Aldrich; St. Louis, Mo.; USA), von Willebrand factor (proteinsynthesized by ECs; baboon explants: Dako M0616, 1:25 dilution; canineexplants: not stained), CD3 (part of the T-cell receptor complex onmature T-lymphocytes; Abcam ab699, 1:25 dilution (Abcam, Cambridge,UK)), CD20 (protein expressed on the surface of mature B cells; Abcamab9475, 1:25 dilution), collagen types I and III (Novus NB600-1408 andNB600-594, 1:200 dilution for both), fibronectin (Novus NB110-1635, 1:50dilution), and vitronectin (Novus NB110-57649, 1:200 dilution) (NovusBiologicals, Cambridge, UK) with either a fluorescent or3,3′-diaminobenzidine (DAB) stain. Alizarin-stained sections wereevaluated to confirm absence of calcification (7±1 sections/animal, n=11animals). Immunogenicity was further evaluated by observation ofH&E-stained sections (11±2 sections/animal, n=12 animals) andimmunostaining for CD3 and CD20 (3±1 sections/animal, n=2 animals).Neointimal thicknesses of native vessels at anastomoses were calculatedas the total area of neointima divided by the length of the underlyingtissue (Lumsden, et al., J Vasc Surg 24, 825-833 (1996)). Amicroscope-mounted camera and image analysis software were used formeasurements.

Statistical Analysis:

Statistical analyses were performed with a Student's two-sample t-test,assuming unequal variances, for analyses with two groups. One-way ANOVAwas used to determine significant differences between three or moregroups, with Tukey's post-hoc comparison. Linear regression wasperformed to assess whether construct suture strength plotted as afunction of cell donor age had a slope significantly different from 0.Two-sided P values less than 0.05 indicated statistical significance.Numeric values are presented as the mean+/−standard error of the mean.Reported ‘n’ represents the number of individual cultured constructstested (not repeat segments from the same graft), and is reported inparenthesis in the tables

Example 3

Generation of Extracellular Matrix Protein Constructs from AllogeneicCells and Decellularization:

To produce extracellular matrix protein constructs (3-6 mm in diameter),allogeneic smooth muscle cells (SMCs) obtained from cadaveric donors arecultured on rapidly degradable poly-glycolic acid (PGA) tubularscaffolds in a bioreactor that delivers cyclic radial strain (Niklason,et al., Science 284, 489-493 (1999)). During the culture period, SMCssecrete extracellular matrix proteins, predominantly collagen, to formbiosynthetic vascular tissue (Niklason, et al., Science 284, 489-493(1999)), and the PGA degrades. At the end of the culture period, theresultant tissue is decellularized with detergents, leaving only thesecreted collagenous matrix (Dahl, et al., Cell Transplantation 12,659-666 (2003)). The decellularization process kills cells, and removesantigenic, allogeneic cells from the construct, thereby allowing the useof banked allogeneic cells to produce extracellular matrix proteinconstructs that are non-immunogenic and can be used in any recipient.These extracellular matrix protein constructs can be stored in astandard phosphate buffered saline (PBS) at 4° C. and immediatelyavailable for arteriovenous access creation (6 mm diameter), or forseeding with autologous ECs to reduce the risk of thrombosis associatedwith small diameter vascular grafting in the peripheral or coronarysettings (3-4 mm graft diameters) (McAllister, et al., Lancet 373,1440-1446 (2009); Kaushal, et al., Nat Med 7, 1035-1040 (2001); Zilla,et al., Semin Vasc Surg 12, 52-63 (1999)).

Example 4

Strength and Stability of Decellularized Human Extracellular MatrixProtein Constructs:

Thirty seven decellularized extracellular matrix protein constructs (6mm diameter, 23 cm length) were produced using cells from 19 humandonors, in order to assess the mechanical consistency of constructsproduced from different donors (Table 2). Suture strength was measuredusing 6-0 Prolene® 2 mm from the edge of every construct. Burst pressurewas tested intermittently on randomly selected constructs by inflating 2cm of tubular construct with saline until rupture. Suture strengths(Table 2) did not change significantly with donor age (P=0.26; ages 17to 49), with male versus female donor origin (P=0.52), or with use ofsingle donor versus pooled donor populations (2-6 donors per pool) forgraft culture (P=0.42). A group of extracellular matrix proteinconstructs were randomly selected and stored for 12 months.Extracellular matrix protein constructs retained their strength, withoutsignificant changes in suture strength, burst pressure, or compliance(P=0.97, P=0.18, and P=0.48, respectively) after 12 months of storage at4° C. in phosphate buffered saline (PBS), and were within the rangesreported for native human vasculature (Table 3). Extracellular matrixprotein constructs contained residual PGA fragments (1.1±0.1% ofcross-sectional area of extracellular matrix protein constructshistological sections prior to storage), which did not degrade furtherduring storage at 4° C. (1.0±0.1% after 9 months of storage, P=0.54).

Table 2 shows donor data and suture strengths for 6 mm-diameterdecellularized human extracellular matrix protein constructs. All dataare presented as Mean±SEM (number of distinct grafts tested).

TABLE 2 Human Donor Donor g Suture Donor Age Sex Diabetic SmokerHypertension Other Diseases Strength 1 17 F No No No Mitral Valve 250(1) Prolapse; Asthma 2 19 M No No No None 233 ± 20 (4) 3 25 F No No NoNone 130 ± 20 (2) 4 33 F No No No None  80 (1) 5 34 F No Yes No None 223± 27 (4) 6 45 F No Yes Yes Asthma 120 ± 10 (2) 7 46 M No Yes Yes None110 (1) 8 46 M Yes Yes No None 115 ± 5 (2)  9 46 M Yes Yes Yes KidneyFailure 135 ± 35 (2) 10  47 M No No No Gastroesophageal 275 ± 42 (4)Reflux Disease 11  47 M No Yes No None 120 ± 30 (2) Pool of 46 M Yes YesYes Kidney Failure 155 ± 35 (2) Donors 9 43 M No Yes Yes None and 12Pool of 17 F No No No Mitral Valve 140 (1) Donors 1, Prolapse; Asthma 4,5, 7, 8 33 F No No No None and 10 34 F No Yes No None 46 M No Yes YesNone 46 M Yes Yes No None 47 M No No No Gastroesophageal Reflux DiseasePool of 18 M No No No None 146 ± 5 (7)  Donors 27 F No No No None 13,14, 43 M No No Yes Herpes and 15 Pool of 17 F No No No None 265 ± 25 (2)Donors 27 M No Yes No None 16, 17, 18, 47 M No Yes Yes Arthritis and 1949 F No Yes No None

Table 3 shows mechanical properties of extracellular matrix proteinconstructs before and after 12 months of storage, and nativecomparators.

TABLE 3 g Suture mmHg Burst % Compliance Strength Pressure per 100 mmHg6 mm diameter  178 ± 11 (37)  3337 ± 343 (10)  3.3 ± 0.8 (10) humanconstructs 6 mm diameter 170 ± 22 (9) 2651 ± 329 (5) 2.5 ± 0.8 (5) humanconstructs stored 12 months Human saphenous 196 ± 29 (7) 1599 ± 877 (7)0.7-1.5 vein Human internal 138 ± 50 (6)  3196 ± 1264 (16) 11.5 ± 3.9(7)  mammary artery

Table 4 shows the minimum and maximum wall thickness and Table 5 showsthe beta-actin, lipid and hydroxyproline content of the decellularizedextracellular matrix protein constructs of the present invention ascompared to cellular based constructs or fresh native tissue.

TABLE 4 Hydroxy- Min Wall Max Wall Beta Actin Lipid proline ThicknessThickness (ng/cm graft (% dry (mg/g dry (um) (um) length) weight)weight) 391 ± 586 ± 41 ± 1.1 ± 60 ± 22 (23) 33 (23) 4 (22) 0.1 (23) 2(14)

Previous reports on engineered tissues made from human cells haveyielded significantly lower suture strengths (59 g), significantly lowerburst pressures (59-108 mmHg), significantly lower average wallthicknesses (181 micrometers), and significantly lower hydroxyprolinecontent (5-16 mg/g dry weight) (Poh, et al., Lancet 365, 2122-2124(2005); McKee, et al., EMBO Rep 4, 633-638 (2003)). This superiority ofthe current constructs as compared to those of Poh et al may be due tothe use of lower-passage human cells, as well as the specific mediumcomposition that was utilized during bioreactor culture, both of whichmay have contributed to superior cell growth, collagenous matrixproduction, and hence improved mechanical properties.

TABLE 5 Beta Actin (ng/cm Lipid (% dry graft length) weight) Fresh 1958± 479 (7) 1.9 ± 0.2 (6)  Decell  41 ± 4 (22) 1.1 ± 0.1 (23)

The constructs of the present invention are devoid of living cells. Betaactin and Lipid contents are reduced via decellularization. In contrast,a devitalization approach that kills cells, but does not remove thecellular remnants, likely retains concentrations of beta actin and lipidthat resemble concentrations of a fresh tissue.

Example 5

Decellularized Human Extracellular Matrix Protein Constructs in anArteriovenous Model:

To assess the function of 6 mm extracellular matrix protein constructs,nine extracellular matrix protein constructs grown from human cells (6mm diameter, 12.5±1.1 cm length) were implanted into baboons asarteriovenous conduits (FIG. 2A) and were observed for 1-6 months (Table6). One animal was excluded after pulling open the surgical incisionsite, exposing the construct and creating a wound infection. Noinfection was observed in the 8 remaining animals. Duplex ultrasoundmeasurements of extracellular matrix protein constructs at weeks 0, 2,4, 12, and 24 (Table 7) showed no change in diameter (P=0.28), no changein wall thickness (P=0.93), and an increase in flow rate between weeks 0and 2 (P<0.01). Flow through extracellular matrix protein constructs(Table 7) was sufficient for hemodialysis (>300 mL/min (B. Dixon, KidneyInt 70, 1413-1422 (2006)). Extracellular matrix protein constructs wereaccessed initially at 4 weeks (FIG. 2B), which is a clinically relevanttime for first access to allow for integration and remodeling ofhemodialysis grafts, and then at 3 and 6 months. Of the eightarteriovenous extracellular matrix protein constructs, 2/2 were patentat 1-month, 2/3 were patent at 3-month explant, and 3/3 were patent at6-month explant (FIG. 2C). Only one construct showed thrombosis at 3months, likely due to technical difficulties with access, which requiredprolonged manual pressure that led to clotting. Hence, the patency ofthe arteriovenous 6 mm extracellular matrix protein constructs in thebaboon was 88% (⅞). No aneurysmal dilatation and no calcification wereobserved in any construct. Furthermore, constructs did not exhibitsubstantial intimal hyperplasia. Anastomotic neointimal hyperplasia at 6months (luminal hyperplasia thickness of 0.11±0.05 mm) was less thanthat reported at 1 month for PTFE arterial bypass grafts in a baboonmodel (0.25±0.09 mm) (Lumsden, et al., J Vasc Surg 24, 825-833 (1996)).

Venous intimal hyperplasia is minimal at 3 months for extracellularmatrix protein constructs in a baboon model. The venous anastomosis ofan extracellular matrix protein construct is shown in FIG. 8A, with theconstruct (G) on the left and the vein (V) on the right. Only smallpatches of intimal hyperplasia are visible on the vein, and thesepatches are circled. In contrast, FIG. 8B shows that the vein adjacentto a PTFE graft has substantially more intimal hyperplasia (see circledregion as an example of the thickness of intimal hyperplasia) in ahuman. Further, the extent of intimal hyperplasia in the graft sectionis also minimal for the extracellular matrix protein construct andsubstantial for the PTFE graft (Prichard et al., An early study on themechanisms that allow tissue-engineered vascular grafts to resistintimal hyperplasia. J Cardiovasc Transl Res 4 (5):674-682, 2011).

Table 6 shows a summary of implanted extracellular matrix proteinconstructs.

TABLE 6 Graft Autol- Inner ogous Graft Access Aneurys- Diameter EC'sTime Points mal Di- (mm) Seeded (Months) latation Patent HumanConstructs in a Baboon Arteriovenous Model 1 Month 6 No No No Yes Access1 Month 6 No 1 No Yes 3 Months 6 No 1 No No 3 Months 6 No 1, 3 No Yes 3Months 6 No 1, 3 No Yes 6 Months 6 No 1, 3, 6 No Yes 6 Months 6 No 1, 3,6 No Yes 6 Months 6 No 1, 3, 6 No Yes Excluded 6 No Excluded ExcludedExcluded Canine Constructs in a Canine Carotid Artery Bypass Model 1Week 3 Yes NA No No 1 Month 3 Yes NA No Yes 12 Months 3 Yes NA No Yes 12Months 4 Yes NA No Yes Excluded 3 Yes Excluded Excluded Excluded CanineConstructs in a Canine Coronary Artery Bypass Model 1 Week 3 Yes NA NoYes 1 Month 3 Yes NA No Yes Excluded 3 Yes Excluded Excluded Excluded

Table 7 shows duplex ultrasound measurements of extracellular matrixprotein constructs placed as arteriovenous grafts in baboons.

TABLE 7 Week 0 Week 2 Week 4 Week 12 Week 24 Diameter (mm) 5.8 ± 0.2 (7)6.3 ± 0.3 (7) 6.7 ± 0.3 (7) 6.8 ± 0.6 (5) 6.3 ± 0.2 (3) Wall thickness(mm) 1.0 ± 0.1 (7) 0.9 ± 0.1 (7) 1.0 ± 0.1 (7) 1.1 ± 0.2 (5) 1.0 ± 0.1(3) Flow rate (ml/min) 764 ± 216 (7) 2278 ± 430 (7)  1464 ± 124 (7) 1559 ± 379 (5)  1572 ± 301 (3) 

Example 6

Extracellular Matrix Protein Constructs in Small Diameter Peripheral andCoronary Arterial Bypass Models:

The function of small diameter (3-4 mm) extracellular matrix proteinconstructs was evaluated in canine models of peripheral and coronaryartery bypass. Canine extracellular matrix protein constructs wereproduced from allogeneic canine cells, decellularized, and luminallyseeded with autologous ECs from the intended recipient. Attached ECswere elongated and aligned within the lumens of extracellular matrixprotein constructs, but complete EC coverage was never achieved. Rather,EC coverage varied widely between constructs, with a coverage range of0-60% (14±8%) on sections sampled from constructs prior to implant. Ingeneral, canine constructs were less strong than human constructs,although still suitable for implantation (burst pressures were 1618±67mmHg for 3 mm canine grafts; n=39).

Five endothelialized canine extracellular matrix protein constructs (3-5cm in length) were implanted as carotid artery bypass grafts, withfollow-up times of 1 week to 12 months (FIG. 2D). One animal wasexcluded after dying acutely with a patent graft. One graft occluded at1 week. All other constructs remained patent, including two constructsthat were followed for one year (Table 6). A representative angiogram atone year (FIG. 2E) demonstrated excellent long-term patency. No stenosisor dilatation was observed in implanted constructs, and no intimalhyperplasia was observed at anastomoses.

Three endothelialized canine extracellular matrix protein constructs(7-10 cm in length) were also implanted into the left anteriordescending coronary artery of dogs (FIG. 2F) and followed for up to 1month (Table 6). One animal died the day after implantation with apatent construct and was excluded from the study. All coronary arterybypass constructs were patent at 1 week and 1 month explants (FIG. 2G).For all small-diameter canine extracellular matrix protein constructs (atotal of 6 in the carotid and coronary circulations), primary patencywas 83% (⅚)

Example 7

Remodeling of Extracellular Matrix Protein Constructs In Vivo:

Prior to implant, extracellular matrix protein constructs were smoothand uniform (FIG. 3A). Histological evaluation (FIG. 3B), as well as DNAquantification (0.74±0.10 μg DNA/mg dry tissue weight), demonstratedthat the extent of decellularization of extracellular matrix proteinconstructs was similar to that of other decellularized scaffolds usedclinically (Derwin, et al., J Bone Joint Surg Am 88, 2665-2672 (2006)).The extracellular matrix of the constructs contained collagen types Iand III, which are the most prevalent types in native vasculature, aswell as fibronectin and vitronectin, all with primarily circumferentialalignment (FIG. 3).

After implantation into baboons and canines, extracellular matrixprotein constructs showed considerable remodeling. For all grafts, grossanalysis at explant revealed a smooth inner graft tissue surface withformation of a loose fibrous outer “adventitial” tissue layer (FIG. 4A).Extracellular matrix protein constructs demonstrated a notable lack ofconstrictive fibrotic tissue surrounding grafts at explant (FIG. 4A).Constructs integrated well with the native vasculature at anastomoticsites (FIG. 4B).

Extracellular matrix protein constructs remodeled to becomecompositionally more similar to native artery after implantation. Within3 months after implant, elastin formed in anastomotic sections of graftsexplanted from baboons (FIG. 4C) in regions containing the highestdensity of infiltrated host cells (FIG. 4D). No elastin was observedmidgraft in any explanted extracellular matrix protein constructs.Alpha-smooth muscle actin positive cells, which could be either SMCs ormyofibroblasts, densely populated the full thickness of extracellularmatrix protein constructs near anastomotic sites (FIG. 4E), suggestingmigration from adjacent native vasculature. Actin-positive cellsappeared to infiltrate transmurally from the adventitial-like tissuelayer into extracellular matrix protein constructs in midgraft regions,starting at 6 months in the baboon model (FIG. 4F).

In the canine model, α-smooth muscle actin positive cells began toinfiltrate midgraft sections transmurally by 1 month (FIG. 4G) and wereobserved throughout the midgraft wall by 1 year (FIG. 4H). Host cellinfiltration into midgraft extracellular matrix protein construct wallswas more rapid in the canine model, possibly because shorter grafts wereplaced in the canines or because of differences in species. In bothmodels, there were fewer cells within the extracellular matrix proteinconstruct walls in midgraft sections (FIG. 4F, Fig. G, Fig. H) thansections near the anastomoses (FIG. 4D, Fig. E). Von Willebrand factor(an EC marker) positive cells were observed on luminal surfaces ofextracellular matrix protein constructs both near anastomotic sites andmidgraft in both canine grafts (which were endothelialized prior toimplantation) and baboon grafts (which were not) (FIG. 4I). ECs may havemigrated from anastomosed vascular tissue, migrated transmurally fromsurrounding tissue (Zilla et al., Biomaterials 28, 5009-5027 (2007)), ororiginated from circulating progenitor cells (Asahara, et al., Science275, 964-967 (1997)).

In the baboon study, midgraft extracellular matrix protein constructsegments were saved for mechanical testing and collagen analysis atexplant. Explanted extracellular matrix protein constructs displayedincreased suture strength (276±28 g, n=8, P=0.01), but no significantchanges in burst pressure (3646±582 mmHg, n=4, P=0.67) or compliance(3.4±2.3% per 100 mmHg, n=4, P=0.70) compared to the pre-implant valuesreported in Table 3. Thus, extracellular matrix protein constructs weremechanically robust without complete infiltration of cells into midgraftsections (FIG. 4F) or elastin in midgraft sections. No significantchanges in collagen density were observed between extracellular matrixprotein constructs pre-implant (57±5%, n=8), extracellular matrixprotein constructs at explant (46±5%, n=7), inflow axillary artery(46±5%, n=7), or control axillary artery explanted from the non-implantarm (42±3%, n=7; P=0.07).

Extracellular matrix protein constructs were not immunogenic. Injectionsof homogenized extracellular matrix protein construct, and PBS as anegative control, were placed intradermally in every baboon at the timeof graft implant and again 4 weeks post implantation (FIG. 5A). Theabsence of visible induration or redness at all injection sitesindicated that recipients were not sensitized to graft material.Immunogenicity of grafts was also assessed by sampling blood frombaboons with implanted extracellular matrix protein constructs, andmeasuring in vitro proliferation of T-cells exposed to PTFE grafts(negative control) or extracellular matrix protein constructs (FIG. 5B).Immunostaining of dense cellular regions (FIG. 5C) showed only sparsepopulations of CD3 or CD20 positive cells (FIG. 5D, Fig. E), which wereoften undetectable in midgraft sections. Foreign body giant cells werenot observed in any explanted extracellular matrix protein construct.Finally, calcification, which is commonly observed in xenogenic orelastin-containing vascular grafts (Hilbert, et al., J Biomed Mater ResA 69, 197-204 (2004); Hopkins, et al., J Thorac Cardiovasc Surg 137,907-913, 913e901-904 (2009)), was not observed in any extracellularmatrix protein construct in any model (FIG. 5F).

Example 8

Extracellular matrix protein constructs were generated by culturinghuman cadaveric donor cells or canine cells on a degradable PGA scaffoldto support synthesis of a collagenous extracellular matrix. Antigeniccellular material was removed via a detergent-based decellularizationstep to render tissues non-immunogenic. The extracellular matrix proteinconstructs contained minimal PGA fragments and retained mechanicalproperties similar to native vessels after 12 months of storage inbuffer at 4° C. Function of 6 mm diameter human extracellular matrixprotein constructs was demonstrated in a baboon arteriovenous model.Small diameter (3-4 mm) canine extracellular matrix protein constructswere luminally seeded with ECs, and implanted in canine models ofperipheral and coronary bypass. Extracellular matrix protein constructsintegrated well with native vasculature at anastomotic sites andresisted intimal hyperplasia. Infiltration of α-smooth muscle actinpositive cells, ECs on graft lumens, and elastin formation nearanastomoses was observed. Long-term patency was demonstrated for up toone year.

One approach of using allogeneic human cells to produce extracellularmatrix protein constructs allows one human donor to provide grafts fordozens of patients (FIG. 6). This approach differs significantly fromthe one-donor-to-one-recipient model, which pertains to autologoustissue engineering and to cadaveric human or animal blood vessels. Onehuman donor provides a cell bank large enough to produce 37 largediameter (6 mm internal diameter) extracellular matrix proteinconstructs or 74 small diameter (3 mm ID) extracellular matrix proteinconstructs. Pooling cells from multiple donors allows for the generationof large cell banks, which in turn makes possible the manufacture ofmany extracellular matrix protein constructs per cell bank (i.e.,200-500 units). This offers greater economies of scale than completelyautologous tissue engineering approaches. Further, use of allogeneiccells, combined with decellularization and simple storage methods,allows the culture period for graft production to be moved “off-line.”Therefore, patients have no waiting period for graft production sincethe grafts have already been created and stored. The ability to storegrafts is an important step in making extracellular matrix proteinconstructs available to the patient immediately at their time of need,as opposed to custom made grafts for each patient. This is an importantdeparture from cell-containing products, which generally cannot bestored long term without specialized cryopreservation equipment andlaborious thawing procedures (Pascual, et al., Ann Vasc Surg 15, 619-627(2001)).

Example 9

A porcine model was used to evaluate the conduit for urinary diversion.Both ureters were anastomosed to the conduit in Wallace fashion. Aurinary diversion stent was placed in each ureter to prevent both earlyanastomotic leakage and stricture, during the normal postoperative phaseof ureter swelling. A plug of skin and subcutaneous tissue was removedin order to accommodate the conduit. A cruciate incision was created inthe fascia, the muscle was split, and a cruciate incision was created inthe posterior rectus sheath. Then, the conduit was brought through theabdominal wall, and secured to the skin and subcutaneous tissue withsuture. A skin barrier and ostomy pouch was adhered to the skinsurrounding the stoma for urine collection. Tunneling in theretroperitoneal plane keeps the graft out of the abdominal cavity, whichminimizes risk of forming adhesions between the graft and otherabdominal tissues. This is important because adhesion formation is areal clinical problem. Tunneling the graft in the retroperitoneal plane,the anastomoses with the ureters, and the anastomosis at with the skinat the stoma site, all provide exposure to a source of vascularization,which may aid in resistance of infection. Exposure to the peritoneum,skin, ureters, and urine also may provide a source of cells to populatethe conduit. FIGS. 9A-9E show the usage of the extracellular matrixprotein constructs of the present invention as urinary conduits. Theresults show that the conduits tolerate chronic exposure to urine.Concentrated human urine was collected, and pumped through a segment ofthe urinary conduit graft for 4 weeks at 37° C. After 4 weeks of urinecirculation, the graft resisted active diffusion of urine through thegraft wall. This lack of significant diffusion across the wall of theconduit is demonstrated by the observation that concentrated urine inthe bottle was darker than the liquid external to the graft in the flowloop (this liquid was phosphate buffered saline, PBS). In contrast,ileal conduits, which are the current gold standard graft material forurinary diversions, actively absorb their contents.

Table 8 shows suture strength of the urinary conduit before and after 4weeks of urine exposure.

TABLE 8 Suture Pullout Strength (g) Graft prior to urine exposure 280 ±35 Graft in urine for 4 weeks 275 ± 15

Other Embodiments

While the invention has been described in conjunction with the detaileddescription thereof, the foregoing description is intended to illustrateand not limit the scope of the invention, which is defined by the scopeof the appended claims. Other aspects, advantages, and modifications arewithin the scope of the following claims.

We claim:
 1. A construct comprising a tubular biodegradable non-wovenpolyglycolic acid scaffold comprising non-biodegradable polyethyleneterephthalate supports at each end of the tubular biodegradablepolyglycolic acid scaffold, wherein the density of the polyglycolic acidis 45 mg/cc to 75 mg/cc and said density is uniform across the entiretubular scaffold, wherein the thickness of the polyglycolic acidscaffold is 0.8 to 1.2 mm, wherein the polyglycolic acid scaffoldcomprises polyglycolic acid fibers and the thickness of the polyglycolicacid fibers within the polyglycolic acid scaffold is 5 to 20 μm, whereinthe porosity of the polyethylene terephthalate is >200 cc/min/cm² andthe supports permit the attachment and growth of cells, wherein thelength of the tubular biodegradable polyglycolic acid scaffold is atleast 10 cm, and wherein the construct is substantially free of heavymetal contaminants.
 2. The construct of claim 1, wherein the length ofthe tubular biodegradable polyglycolic acid scaffold is 10 cm to 100 cm.3. The construct of claim 1, wherein the inner diameter of the tubularbiodegradable polyglycolic acid scaffold is 3 mm to 6 mm.
 4. Theconstruct of claim 1, wherein the construct comprises trace amounts ofheavy metal contaminants selected from the group consisting of:aluminum, barium, calcium, iodine, lanthanum, magnesium, nickel,potassium and zinc.
 5. The construct of claim 1, wherein the constructfurther comprises extracellular matrix proteins.
 6. The construct ofclaim 5, wherein the thickness of the extracellular matrix proteins isgreater than 200 micrometers at the thinnest portion of the matrix. 7.The construct of claim 1, wherein the construct comprises a uniformlyentangled seam having a density that matches the overall density of thetubular construct.